1. Field
The present disclosure generally relates to Computed Tomography (CT) imaging. In particular, embodiments herein relate to an apparatus and method for controlling movement of detectors mounted on a ring in a CT system.
2. Background
Radiographic imaging, in its simplest expression, is an X-ray beam traversing an object and a detector relating the overall attenuation per ray. The attenuation is derived from a comparison of the same ray with and without the presence of the object. From this conceptual definition, several steps are required to properly construct an image. For instance, the finite size of the X-ray generator, the nature and shape of the filter blocking the very low energy X-ray from the generator, the details of the geometry and characteristics of the detector, and the capacity of the acquisition system are all elements that affect how the actual reconstruction is performed.
In one of many possible geometries, the X-ray source on top of the graph shown in FIG. 1 is emitting an X-ray beam forming a fan, traversing the object. While a wide range of values can exist, typically, the distance “C” is around 100 cm, “B” is around 60 cm, and “A” is around 40 cm. The principle of tomography requires that each point of the object is traversed by a collection of rays covering at least 180 degrees. Thus, the entire X-ray generator and detector assembly will rotate around an object. Mathematical considerations show that the tomographic conditions are met when a scan of 180 degrees plus the fan angle is performed.
Conventional X-ray detectors integrate the total electrical current produced in a radiation sensor, and disregard the amplitude information from individual photon detection events. Since the charge amplitude from each event is proportional to the photon's detected energy, this acquisition provides no information about the energy of individual photons, and is thus unable to capture the energy dependence of the attenuation coefficient in the object.
On the other hand, semiconductor X-ray detectors that are capable of single photon counting and individual pulse height analysis may be used. These X-ray detectors are made possible by the availability of fast semiconductor radiation sensors materials with room temperature operation and good energy resolution, combined with application-specific integrated circuits (ASICs) suitable for multi-pixel parallel readout and fast counting.
When operating such a photon counting X-ray detector, a high bias voltage is applied across the sensor crystal such that the electron-hole pairs generated from the radiation interaction are rapidly swept toward the collecting electrodes. Each radiation interaction event results in a pulse sent to the readout electronics, which undergoes pulse height analysis and is counted.
One major advantage of such photon counting detectors is that, when combined with pulse height analysis readout, spectral information can be obtained about the x-ray beams passing through the object and then the attenuation coefficient at each energy in the object can be reconstructed. Conventional CT measures the attenuation at one average energy only, while in reality, the attenuation coefficient strongly depends on the photon energy. In contrast, with pulse height analysis, a system is able to categorize the incident X-ray photons into several energy bins based on their detected energy. This spectral information can effectively improve material discrimination and target contrast, all of which can be traded for a dose reduction to, for example, a patient.
In fourth-generation spectral CT, the photon counting detectors (PCDs) are located on a ring and fixed on the gantry. As the X-ray source rotates, a PCD will “see” the X-ray source from a different angle. In other words, the X-ray beam will irradiate the PCD surface from different angles. On different angles, the PCDs may have different count response and energy response. This angular variation may complicate the detector response function and then data domain decomposition. Additionally, this also makes the fixed PCDs susceptible to unwanted scattered radiation.
In third-generation CT, standard practice is to use an anti-scatter grid, which is placed between the patient and the detector. However, in fourth-generation CT, the angular variation of the incident beam makes use of the anti-scatter grid challenging. Some fourth-generation CT scanners (e.g., single slice scanners) use a fixed detector array just outside the incident beam to detect scattered photons. Such arrangement helps estimate scatter for subsequent correction, but does not solve the count and energy response variation problem.